Magnetic resonance imaging gradient coil

ABSTRACT

A magnetic gradient coil (110) for a magnetic resonance imaging system (100, 200) is actively shielded. The magnetic gradient coil is operable for generating a magnetic field (504). The magnetic field has a cylindrical axis of symmetry (130). The gradient coil has a length (132) parallel with the cylindrical axis of symmetry. The magnetic gradient coil has an outer surface (134). The magnetic field includes an external magnetic field outside of the outer surface. The external magnetic field has at least four reduced field regions (136, 138, 140, 142) along the length where the modulus of the magnetic field is less than the average of the modulus of the magnetic field along the length.

CROSS REFERENCE TO RELATED APPLICATIONS

This application is a U.S. national phase application of InternationalApplication No. PCT/EP2014/062528, filed on Jun. 16, 2014, which claimsthe benefit of EP Application Serial No. 13172199.5 filed on Jun. 17,2013 and is incorporated herein by reference.

TECHNICAL FIELD

The invention relates to magnetic resonance imaging, in particular tothe design and construction of magnetic gradient coils.

BACKGROUND OF THE INVENTION

A magnetic field is used in Magnetic Resonance Imaging to align thenuclear spins of atoms as part of the procedure for producing imageswithin the body of a patient. This magnetic field is referred to as theB0 field. During an MRI scan, Radio Frequency (RF) pulses generated by atransmitter or amplifier and an antenna cause perturbations to the localmagnetic field and can be used to manipulate the orientation of thenuclear spins relative to the B0 field. Spatial encoding of the magneticspins may be accomplished by using so called gradient coils, which areused to superimpose a magnetic field gradient upon the B0 magneticfield. RF signals emitted by the nuclear spins are detected by areceiver coil, and these RF signals are used to construct the MRIimages.

The magnets used to generate the B0 field typically use superconductivecoils. The magnetic field generated by the gradient coils can cause eddycurrents within the superconductive coils. These eddy currents can beavoided or reduced by using gradient coils with active shielding. U.S.Pat. No. 4,733,189 discloses an active shield about the gradient-formingcomponents of a magnetic resonance imaging system. The actively shieldedgradient magnetic field coil known from JP2008-229360 has a fielddistribution expanding radially and having three longitudinal regions inwhich a leakage gradient field expands.

SUMMARY OF THE INVENTION

The invention provides for a magnetic gradient coil, a magnet assembly,a magnetic resonance imaging system and a method and design of amagnetic gradient coil in the independent claims. Embodiments are givenin the dependent claims.

As will be appreciated by one skilled in the art, aspects of the presentinvention may be embodied as an apparatus, method or computer programproduct. Accordingly, aspects of the present invention may take the formof an entirely hardware embodiment, an entirely software embodiment(including firmware, resident software, micro-code, and etc.) or anembodiment combining software and hardware aspects that may allgenerally be referred to herein as a “circuit,” “module” or “system.”Furthermore, aspects of the present invention may take the form of acomputer program product embodied in one or more computer readablemedium(s) having computer executable code embodied thereon.

Any combination of one or more computer readable medium(s) may beutilized. The computer readable medium may be a computer readable signalmedium or a computer readable storage medium. A ‘computer-readablestorage medium’ as used herein encompasses any tangible storage mediumwhich may store instructions which are executable by a processor of acomputing device. The computer-readable storage medium may be referredto as a computer-readable non-transitory storage medium. Thecomputer-readable storage medium may also be referred to as a tangiblecomputer readable medium. In some embodiments, a computer-readablestorage medium may also be able to store data which is able to beaccessed by the processor of the computing device. Examples ofcomputer-readable storage media include, but are not limited to: afloppy disk, a magnetic hard disk drive, a solid state hard disk, flashmemory, a USB thumb drive, Random Access Memory (RAM), Read Only Memory(ROM), an optical disk, a magneto-optical disk, and the register file ofthe processor. Examples of optical disks include Compact Disks (CD) andDigital Versatile Disks (DVD), for example CD-ROM, CD-RW, CD-R, DVD-ROM,DVD-RW, or DVD-R disks. The term computer readable-storage medium alsorefers to various types of recording media capable of being accessed bythe computer device via a network or communication link. For example adata may be retrieved over a modem, over the internet, or over a localarea network. Computer executable code embodied on a computer readablemedium may be transmitted using any appropriate medium, including butnot limited to wireless, wire line, optical fiber cable, RF, etc., orany suitable combination of the foregoing.

A computer readable signal medium may include a propagated data signalwith computer executable code embodied therein, for example, in basebandor as part of a carrier wave. Such a propagated signal may take any of avariety of forms, including, but not limited to, electro-magnetic,optical, or any suitable combination thereof. A computer readable signalmedium may be any computer readable medium that is not a computerreadable storage medium and that can communicate, propagate, ortransport a program for use by or in connection with an instructionexecution system, apparatus, or device.

‘Computer memory’ or ‘memory’ is an example of a computer-readablestorage medium. Computer memory is any memory which is directlyaccessible to a processor. ‘Computer storage’ or ‘storage’ is a furtherexample of a computer-readable storage medium. Computer storage is anynon-volatile computer-readable storage medium. In some embodimentscomputer storage may also be computer memory or vice versa.

A ‘processor’ as used herein encompasses an electronic component whichis able to execute a program or machine executable instruction orcomputer executable code. References to the computing device comprising“a processor” should be interpreted as possibly containing more than oneprocessor or processing core. The processor may for instance be amulti-core processor. A processor may also refer to a collection ofprocessors within a single computer system or distributed amongstmultiple computer systems. The term computing device should also beinterpreted to possibly refer to a collection or network of computingdevices each comprising a processor or processors. The computerexecutable code may be executed by multiple processors that may bewithin the same computing device or which may even be distributed acrossmultiple computing devices.

Computer executable code may comprise machine executable instructions ora program which causes a processor to perform an aspect of the presentinvention. Computer executable code for carrying out operations foraspects of the present invention may be written in any combination ofone or more programming languages, including an object orientedprogramming language such as Java, Smalltalk, C++ or the like andconventional procedural programming languages, such as the “C”programming language or similar programming languages and compiled intomachine executable instructions. In some instances the computerexecutable code may be in the form of a high level language or in apre-compiled form and be used in conjunction with an interpreter whichgenerates the machine executable instructions on the fly.

The computer executable code may execute entirely on the user'scomputer, partly on the user's computer, as a stand-alone softwarepackage, partly on the user's computer and partly on a remote computeror entirely on the remote computer or server. In the latter scenario,the remote computer may be connected to the user's computer through anytype of network, including a local area network (LAN) or a wide areanetwork (WAN), or the connection may be made to an external computer(for example, through the Internet using an Internet Service Provider).

Aspects of the present invention are described with reference toflowchart illustrations and/or block diagrams of methods, apparatus(systems) and computer program products according to embodiments of theinvention. It will be understood that each block or a portion of theblocks of the flowchart, illustrations, and/or block diagrams, can beimplemented by computer program instructions in form of computerexecutable code when applicable. It is further under stood that, whennot mutually exclusive, combinations of blocks in different flowcharts,illustrations, and/or block diagrams may be combined. These computerprogram instructions may be provided to a processor of a general purposecomputer, special purpose computer, or other programmable dataprocessing apparatus to produce a machine, such that the instructions,which execute via the processor of the computer or other programmabledata processing apparatus, create means for implementing thefunctions/acts specified in the flowchart and/or block diagram block orblocks.

These computer program instructions may also be stored in a computerreadable medium that can direct a computer, other programmable dataprocessing apparatus, or other devices to function in a particularmanner, such that the instructions stored in the computer readablemedium produce an article of manufacture including instructions whichimplement the function/act specified in the flowchart and/or blockdiagram block or blocks.

The computer program instructions may also be loaded onto a computer,other programmable data processing apparatus, or other devices to causea series of operational steps to be performed on the computer, otherprogrammable apparatus or other devices to produce a computerimplemented process such that the instructions which execute on thecomputer or other programmable apparatus provide processes forimplementing the functions/acts specified in the flowchart and/or blockdiagram block or blocks.

A ‘user interface’ as used herein is an interface which allows a user oroperator to interact with a computer or computer system. A ‘userinterface’ may also be referred to as a ‘human interface device.’ A userinterface may provide information or data to the operator and/or receiveinformation or data from the operator. A user interface may enable inputfrom an operator to be received by the computer and may provide outputto the user from the computer. In other words, the user interface mayallow an operator to control or manipulate a computer and the interfacemay allow the computer indicate the effects of the operator's control ormanipulation. The display of data or information on a display or agraphical user interface is an example of providing information to anoperator. The receiving of data through a keyboard, mouse, trackball,touchpad, pointing stick, graphics tablet, joystick, gamepad, webcam,headset, gear sticks, steering wheel, pedals, wired glove, dance pad,remote control, and accelerometer are all examples of user interfacecomponents which enable the receiving of information or data from anoperator.

A ‘hardware interface’ as used herein encompasses an interface whichenables the processor of a computer system to interact with and/orcontrol an external computing device and/or apparatus. A hardwareinterface may allow a processor to send control signals or instructionsto an external computing device and/or apparatus. A hardware interfacemay also enable a processor to exchange data with an external computingdevice and/or apparatus. Examples of a hardware interface include, butare not limited to: a universal serial bus, IEEE 1394 port, parallelport, IEEE 1284 port, serial port, RS-232 port, IEEE-488 port, Bluetoothconnection, Wireless local area network connection, TCP/IP connection,Ethernet connection, control voltage interface, MIDI interface, analoginput interface, and digital input interface.

A ‘display’ or ‘display device’ as used herein encompasses an outputdevice or a user interface adapted for displaying images or data. Adisplay may output visual, audio, and or tactile data. Examples of adisplay include, but are not limited to: a computer monitor, atelevision screen, a touch screen, tactile electronic display, Braillescreen, Cathode ray tube (CRT), Storage tube, Bi-stable display,Electronic paper, Vector display, Flat panel display, Vacuum fluorescentdisplay (VF), Light-emitting diode (LED) displays, Electroluminescentdisplay (ELD), Plasma display panels (PDP), Liquid crystal display(LCD), Organic light-emitting diode displays (OLED), a projector, andHead-mounted display.

Magnetic Resonance (MR) data is defined herein as being the recordedmeasurements of radio frequency signals emitted by atomic spins by theantenna of a Magnetic resonance apparatus during a magnetic resonanceimaging scan. Magnetic resonance data is an example of medical imagedata. A Magnetic Resonance Imaging (MRI) image is defined herein asbeing the reconstructed two or three dimensional visualization ofanatomic data contained within the magnetic resonance imaging data. Thisvisualization can be performed using a computer.

In one aspect the invention provides for a magnetic gradient coil for amagnetic resonance imaging system. The magnetic gradient coil isactively shielded. An actively shielded magnetic gradient coil comprisesmultiple layers with windings that are used to tailor the magnetic fieldgenerated by the gradient coil. It is generally desirable to create themagnetic gradient field in the region where imaging is being performedon a subject. The external magnetic field from the gradient coils mayinterfere with the superconducting magnet used for generating the mainor so called B0 magnetic field used for magnetic resonance imaging. Suchmagnetic coupling can result in induced currents (which produce unwantedmagnetic fields), heating and, in extreme cases, even gradual or suddenloss of field of the main field magnet.

The shielding windings of the actively shielded magnetic gradient coilsare used to reduce or eliminate the magnetic field generated by themagnetic gradient coil outside of it. The magnetic gradient coil isoperable for generating a magnetic field. The main field magnet has acylindrical axis of symmetry. The main (B0) field is directed along thisaxis of symmetry. The gradient coils are designed such as to generatedefined gradients in the field component directed along the cylindricalaxis of symmetry; for reasons of conservation of magnetic flux, theother field components will be present as well. It may be in someinstances that the axis of symmetry is also a mechanical axis ofsymmetry with respect to the housing or case of the magnetic gradientcoil; however this does not need to be the case. For instance the innerboundary of the gradient coil assembly need not be cylindrical. It maybe elliptical, prismatic or may be asymmetrical in the up or downdirection. The gradient coil has a length parallel with the cylindricalaxis of symmetry. The length as used herein is simply a direction orpath in space. The magnetic gradient coil has an outer surface. Themagnetic field comprises an external magnetic field outside of the outersurface. The external field has at least four reduced field regionsalong the length where the modulus magnetic field is less than theaverage of the modulus of the magnetic field along the length. In otherwords if one goes outside of the magnetic gradient coil at an exteriorsurface of the gradient coil and measures the magnetic field along adirected path that is parallel with the cylindrical axis of symmetrythere will be at least four regions along the length where the modulusmagnetic field is less than the average of the modulus of the magneticfield along the length.

Conventional magnetic gradient coils are designed such that the magneticfield outside of the magnetic gradient coil is reduced in the entirevolume directly surrounding the outer boundary of the gradient coil.Embodiments of the invention are engineered such that there are reducedfield regions only in discreet locations. These may be lined up oraligned with coils in the main magnet. By looking at the magnetic fieldoutside of the magnetic gradient coil it is straight forward todifferentiate a magnetic gradient coil according to an embodiment N1which is 0. Such a magnetic gradient coil as described herein may haveseveral different advantages. First of all the magnetic field isessentially allowed to expand in between the locations of where coils inthe main magnet could be located. This makes the magnetic gradient coilmore efficient while still reducing unwanted magnetic coupling with themain magnet. This also enables the use of a power supply with reducedpower requirements. The reduction in the power requirements also enablesthe use of a magnetic gradient coil power supply with reduced powerrequirements and/or power consumption.

A magnetic gradient coil as used herein encompasses one or coil forsuperimposing a so called gradient magnetic field on an imaging zone orregion when performing magnetic resonance imaging. The magnetic gradientcoil is used for spatially encoding nuclear spins so that spatiallyresolved images can be reconstructed. References to “a gradient coil” or“one gradient coil” should be interpreted as being one or more or a setof gradient coils. Multiple gradient coils are used in a magneticresonance imaging system to perform spatial encoding inthree-dimensions. For a cylindrical magnet, the axis of symmetry isreferred to as the z-axis. A z-gradient coil performs encoding along thez-axis. Two other gradient coils are typically used to generategradients along an x-axis and y-axis. These other two axes are typicallychosen so that they are orthogonal to each other and to the z-axis. Thegradient coils corresponding to the x-axis and the y-axis are typicallyreferred to as the x-gradient coil and the y-gradient coil respectively.Embodiments include replacing one or more conventional gradient coilswith a gradient coil as described herein. Each of the x-gradient coil,the y-gradient coil, and the z-gradient coils would have its ownseparate layers in the gradient coil or gradient coil assembly.

In one example, the x-gradient coil, the y-gradient coil, and thez-gradient coil are constructed according to an embodiment. In anotherexample, the x-gradient coil and the y-gradient coil are constructedaccording to an embodiment and the z-gradient coil is a conventionalactively shielded gradient coil.

In another embodiment the linear extent of each of the reduced fieldregions as measured along the length is at least 10% of the distancebetween two adjacent reduced field regions. This is beneficial becausethe reduced field region is actually a large volume and is not simply asingle point where the field strength reduces.

In another embodiment the modulus of the external field within any oneof the reduced field regions is at least a factor of 2.5 times smallerthan the average of the modulus of the magnetic field along the length.This embodiment may be beneficial because the reduction in the fieldreduces the coupling between the magnetic gradient coil and the coils inthe main magnet.

In another embodiment the modulus external magnetic field within any oneof the reduced field regions is at least a factor of 5 times smallerthan the average of the modulus magnetic field along the length.

In another embodiment the modulus of the external field within any oneof the reduced field regions is at least a factor of 10 times smallerthan the average of the modulus of the magnetic field along the length.

In another embodiment the modulus of the external field within any oneof the reduced field regions is at least a factor of 20 times smallerthan the average of the modulus of the magnetic field along the length.

In alternative embodiments the modulus of the external field within allof the reduced field regions is at least a factor of 20, 10, 5, or 2.5times smaller than the average of the modulus of the magnetic fieldalong the length.

In another embodiment the gradient coil has an inner conductive layerand an outer conductive layer. The inner conductive layer and outerconductive layer are formed by coils. The inner conductive layercomprises a first set of discreet current loops connected in series andthe outer conductive layer comprises a second set of discreet currentloops connected in series. The first set is connected in series to thesecond set.

In another embodiment the magnetic gradient coil comprises threedistinct gradient coils performing three-dimensional spatial encoding.The reduced field zones of all three gradient coils coincide andcorrespond to the typical positions of primary coils of asuperconducting magnet. These for instance may be the primary coils of asuperconducting whole body magnet used for magnetic resonance imaging.

In another aspect the invention provides for a magnet assembly for amagnetic resonance imaging system comprising the actively shieldedgradient coil according to an embodiment of the invention and a magnet.The magnet is a superconducting magnet with multiple superconductingcoils. The magnet is a cryogen-free magnet (meaning that the sections ofthe magnet are not immersed in a large volume of liquid helium butcooled by thermal contact to an active refrigerator and the spacebetween the sections of the main magnet does not contain large amountsof electrically conducting material such as a tank containing the heliumof a bath cooled magnet).

Cryogen-free magnets may be considered to be magnets that do not uselarge amounts of liquid cryogen.

Cryogen-free magnets use a cryogenic insulation system to be able tocool down the magnet to a temperature of about 4 Kelvin if it usesconventional low temperature superconductors and maybe 30-50K in casehigh temperature superconductors are used. The cryogenic insulationcomprises an outer vacuum container, completely enclosing the magnet,enabling an insulating vacuum around the cold mass with a very lowpressure. The insulation also features at least one radiation shieldinside the vacuum space, which has the function to intercept radiationheat coming from the warm surface of the vacuum container. By thermallyconnecting this radiation shield to the refrigerator, nearly all of thisradiation heat is removed from the system. The shield then provides acold surface facing the cold inner parts of the magnet, which radiatesfar less heat than the room temperature wall. Conventional (bath cooled)magnets typically contain 1000-2000 liters of liquid helium in the tanksurrounding the magnet coils. Cryogen-free magnets can work with ahelium inventory of several liters.

There is a superconducting coil selected from the multiplesuperconducting coils centered coaxially about each reduced fieldregion. In other words the coils of the superconducting coils arealigned or positioned in the reduced field regions. This is beneficialbecause the magnetic field is able to expand between the superconductingcoils of the magnet whereas even though there is reduced couplingbetween the magnetic gradient coil and the multiple coils of thesuperconducting magnet.

If the magnetic resonance imaging magnet for which the magnetic gradientcoil is designed has mirror symmetry, the pattern of reduced fieldregions also has a mirror symmetry or anti-symmetry relative to the Z0plane where Z is an axis of symmetry of the magnetic field of themagnet.

In another embodiment the magnet comprises a warm bore tube and aradiation screen. The radiation screen comprises an inner cylinder ofthe radiation screen between the warm bore tube and the multiplesuperconducting coils.

In another embodiment the warm bore tube and the inner cylinder of theradiation screen are electrically non-conducting or have a much higherelectrical impedance in azimuthal direction than in the axial directionof the cylinder. This embodiment may be beneficial because this wouldenable the magnetic field to penetrate the warm bore tube and the innercylinder of the radiation screen.

In another embodiment the inner cylinder is formed from a dielectric.

In another embodiment the inner cylinder of the radiation screen isformed from a conductive material with slots operable to block eddycurrents generated by the external magnetic field of the magneticgradient coil. This is beneficial because the radiation screen can bemade of a highly thermal conductive material such as a metal, however itwould still allow the magnetic field to pass through it. For example aseries of slits cut parallel to the center of axis of the magnet or thegradient coil would enable this.

In another embodiment the inner cylinder of the radiation screen isformed from a dielectric.

In another embodiment the warm bore tube comprises the magnetic gradientcoil. This embodiment may be beneficial because the gradient coil isused to form the warm bore tube. This eliminates a component.

In another embodiment the magnet is a cryogen-free magnet. Acryogen-free magnet is a magnet without liquid helium. Thesuperconducting windings of the MRI magnet are not enclosed in anelectrically conducting helium tank but surrounded by vacuum. The coilsof the magnet are cooled by means of thermal conduction or with acirculating gas and/or liquid in cooling tubes. These cooling tubes maybe connected to a refrigerator. In another case, the magnet may becooled by means of a helium bath but the helium tank may have anelectrically non-conducting inner cylinder separating the liquid heliumfrom the insulating vacuum in one case.

In another embodiment the gradient coil is operable for producing theexternal magnet field such that it expands between each superconductingcoil. This is beneficial because it reduces the amount of energynecessary for powering the magnetic gradient coil.

In another aspect the invention provides for a magnetic resonanceimaging system comprising a magnet assembly according to an embodimentof the invention.

In another aspect the invention provides for a method of designing amagnetic gradient coil for a magnetic resonance imaging system usingmagnetic design software. Magnetic design software as used hereinencompasses software which accepts as input constraints on the generatedmagnetic field and uses the information to calculate the position ofwindings in order to generate the design of a magnetic gradient coil.The use of such magnetic design software is typically used by thedesigners of magnetic resonance imaging systems. As such the method ofdesigning a magnetic gradient coil is essentially differentiated fromthe known methods by defining the constraints differently than isnormally performed.

The method comprises the step of defining cylindrical surfaces or meshescorresponding to an inner conductive layer and an outer conducting layerof the gradient coil. Optionally this step may include definingconductive flange regions connecting the inner conductive layer and theouter conductive layer. The method further comprises definingconstraints and forcing a magnetic field gradient with a predeterminedlinearity within an imaging volume inside the gradient coil. Themagnetic gradient coil has an outer surface. The method furthercomprises defining constraining limits for an outer magnetic fieldsurrounding the outer surface such that there are at least four reducedfield regions corresponding to the location of the superconducting coilof a superconducting magnet.

Alternatively or including this step may also include defining surfacesacting as passive conducting rings at the locations of the reduced fieldregions in defining constraints for the current induced in these ringsor the dissipation caused by these reduced currents. The previous threesteps define an optimization problem for finding the currentdistribution in inner and outer field generating surfaces satisfying allconstraints and having minimum magnetic stored energy and/or dissipationof the magnetic stored energy. The method further comprises solving theoptimization problem. This for instance would be performed automaticallyby the magnetic design software.

The method further comprises converting a continuous stream functionobtained as a result of optimization into a pattern of discreet currentloops. The stream function is a scalar quantity on a surface, with theproperty that the difference in stream function value between any twopoints is the amount of current passing between these two points. Aniso-contour line plot of the stream function with a certain constantstep-size yields the shapes of discrete windings each carrying a currentequal to this step-size. The theory of stream functions is described inthe doctoral thesis: G. N. Peeren, Stream Function Approach ForDetermining Optimal Surface Currents, Ph.D. Thesis, TechnischeUniversiteit Eindhoven, 2003.

These continuous stream functions can be approximated into a pattern ofdiscreet current loops. This step also comprises connecting thesediscreet current loops in series to define a gradient coil design.

In another embodiment the method further comprises the step ofmanufacturing the gradient coil according to the gradient coil design.This in combination with the other method steps may be considered to bea method of producing a magnetic gradient coil.

The current loop patterns may for instance be produced as wire wound,printed circuit based, or copper sheet punched coil parts.

It is understood that one or more of the aforementioned embodiments ofthe invention may be combined as long as the combined embodiments arenot mutually exclusive.

BRIEF DESCRIPTION OF THE DRAWINGS

In the following preferred embodiments of the invention will bedescribed, by way of example only, and with reference to the drawings inwhich:

FIG. 1 shows an example of a cross-sectional and functional view of amagnetic resonance imaging system;

FIG. 2 shows a further example of a cross-sectional and functional viewof a magnetic resonance imaging system;

FIG. 3 shows a flow diagram which illustrates a method of designing amagnetic gradient coil;

FIG. 4 shows a mesh which defines the magnetic gradient coil in twolayers and the smaller cylinders labeled represent meshes defined forthe coils of the superconducting magnet;

FIG. 5 shows an example of a typical example solution of the relaxedexternal fields;

FIG. 6 shows a graph of the stored energy versus the coil outer radius;

FIG. 7 shows the same plot for the conditions of FIG. 6 except thedissipation of energy is shown instead of the stored energy;

FIG. 8 shows the magnetic field contours in a model of a magnet definedby the superconducting coils;

FIG. 9 Stored energy of partly and fully shielded gradient coils vs.inner radius, for outer radius 415 and 425 mm;

FIG. 10 shows the dissipation of energy versus the inner radius;

FIG. 11 shows the visualization of a coil geometry as was shown in FIG.5;

FIG. 12 shows the conductor pattern of the inner conductive layer; and

FIG. 13 shows the conductor pattern of the outer conductive layer.

DETAILED DESCRIPTION OF THE EMBODIMENTS

Like numbered elements in these figures are either equivalent elementsor perform the same function. Elements which have been discussedpreviously will not necessarily be discussed in later figures if thefunction is equivalent.

FIG. 1 shows a cross-sectional and functional view of a magneticresonance imaging system 100. The magnetic resonance imaging system 100is shown as comprising a magnet 102. The magnet 102 shown in FIG. 1 is acylindrical type superconducting magnet. The magnet 102 has a bore 104through the center of it. However, other magnets are also applicable forembodiments of the invention. The magnet 102 has a cryostat 106. Insidethe cryostat 106 there is a collection of superconducting coils 108. Thesuperconducting coils 108 are not enclosed in an electrically conductinghelium tank. Either the helium tank has an electrically non-conductinginner cylinder, separating the liquid helium from the insulating vacuumor the coils of the magnet are cooled by means of thermal conduction orcirculating gas. As an alternative or additional feature a small amountof gaseous or liquid helium in cooling tubes may be used to cool thesuperconducting coils, wherein the cooling tubes are connected to arefrigerator. Not all details of the magnet are shown in FIG. 1.

Within the bore of the magnet there is a magnetic field gradient coil110 which is used for acquisition of magnetic resonance data tospatially encode objects within an imaging zone 114 of the magnet 102.The magnetic field gradient coil 110 is connected to a magnetic fieldgradient coil power supply 112. The magnetic field gradient coil 110 isintended to be representative. Typically magnetic field gradient coilscontain three separate sets of coils for spatially encoding in threeorthogonal spatial directions. The imaging zone 114 is located in thecentre of the magnet 102.

Adjacent to the imaging zone 114 is a radio frequency coil 116 formanipulating the orientations of magnetic spins within the imaging zone114 and for receiving radio transmissions from spins also within theimaging zone 114. The radio frequency coil 116 is connected to a radiofrequency transceiver 118. The radio frequency coil 116 and radiofrequency transceiver 118 may be replaced by separate transmit andreceive coils and a separate transmitter and receiver. It is understoodthat the radio frequency coil 116 and the radio frequency transceiver118 are representative of the different possibilities.

Within the center of the magnet is also located a subject 120. Thesubject 120 is shown as reposing on a subject support 122. The dashedline 130 represents a cylindrical axis of symmetry. It could be the axisof symmetry for the magnetic gradient coil 110 and/or the magnet 102. Inthis case the axis 130 is symmetric with respect to the housing of boththe magnet 102 and the magnet gradient coil 110. The line 132 is adirected path or a length parallel to the axis 130 that runs along anouter surface of the magnetic gradient coil 134. There are dashed lineswhich divide the magnet 102 and the gradient coil 134 into a series ofsections. These are labeled 135, 136, 137, 138, 139, 140, 141, 142 and143. The regions 136, 138, 140 and 142 are so called reduced fieldregions. The magnetic field along the path 132 is less than the averagealong the entire path 132. The regions 135, 137, 139, 141 and 143 arehigher field regions and the magnetic field from the gradient coil 110is allowed to expand into these regions.

The radio frequency transceiver 118 and the magnetic field gradient coilpower supply 112 are shown as being connected to a hardware interface152 of a computer system 150. The computer system 150 uses a processor154 to control the magnetic resonance imaging system 100.

The computer system 150 shown in FIG. 1 is representative. Multipleprocessors and computer systems may be used to represent thefunctionality illustrated by this single computer system 150. Thecomputer system 150 comprises the hardware interface 152 which allowsthe processor 154 to send and receive messages to components of themagnetic resonance imaging system 100. The processor 154 is alsoconnected to a user interface 156, computer storage 158, and computermemory 160.

The radio-frequency transceiver 118 and the magnetic gradient coil powersupply 112 are connected to a hardware interface 152 of computer system150. The computer storage 158 is shown as containing a pulse sequence.

The computer storage 158 is shown as containing a pulse sequence 162.The pulse sequence 162 is a series of commands or information which maybe used to generate commands for controlling the operation of themagnetic resonance imaging system 100 to acquire magnetic resonancedata. The computer storage is also shown as containing magneticresonance data 164 that was acquired using the pulse sequence 162. Thecomputer storage 158 is also shown as containing a magnetic resonanceimage 166 that was reconstructed from the magnetic resonance data 164.

The computer memory 160 is shown as containing a control module 168. Thecontrol module 168 contains computer-executable code which enables theprocessor 154 to control the operation and function of the magneticresonance imaging system 100. This includes using the pulse sequence 162to acquire the magnetic resonance data 164. The computer memory 160 isfurther shown as containing an image reconstruction module 170. Theimage reconstruction module contains computer-executable code whichenables the processor 154 to perform mathematical functions on themagnetic resonance data 164 to reconstruct the magnetic resonance image166.

FIG. 2 shows a magnetic resonance imaging system 200 similar to thatshown in FIG. 1, however in this example more details of the magnet 102are shown. Also in this FIG. the gradient coil is used as a warm bore206.

The system of superconducting coils is enclosed by an outer vacuumcontainer at room temperature. Inside the outer vacuum container is avacuum region 202. The superconducting coils 108 are positioned insidethe vacuum 202 and are individually cooled by either a liquid or gascooling system connected to a cooling or refrigeration system. Thismagnet 102 has a warm bore 206 that is formed by the magnetic gradientcoil 110. There is a radiation shield 204 between the connecting coils108, the warm bore 206 and the walls of the magnet 102. There is aninner cylinder 208 of the radiation shield between the warm bore 206 andthe superconducting coils 108. Both the warm bore 206 and the innerradius of the radiation shield 208 are operable to allow the magneticfield from the gradient coil 110 to pass through it. This may beaccomplished by either using dielectric materials or particularly in thecase of the inner cylinder of the radiation shield 208 to use aconductive material but cut slots in it to allow the magnetic field toexpand into the higher field regions 135, 136, 139, 141 and 143.

FIG. 3 shows a flow diagram which illustrates a method of designing amagnetic gradient coil. Step 300 is defining a cylindrical surface ormeshes corresponding to an inner conductive layer and an outerconductive layer of the gradient coil and optionally defining conductingflange regions connecting the inner conductive layer and the outerconductive layer. Next in step 302 constraints are defined forcing themagnetic gradient field with a predetermined linearity within an imagingvolume inside the gradient coil. The imaging volume may be the imagingzone. The magnetic gradient coil has an outer surface. Next in step 304constraints limiting a magnetic field surrounding the outer surface aredefined such that there are at least four reduced field regionscorresponding to the location of a superconducting coil of thesuperconducting magnet. This is performed optionally or with definingsurfaces acting as passive conducting rings at the locations of thereduced field regions and defining constraints for the current inducedin these rings or the dissipation caused by these induced currents.

The steps 300, 302 and 304 define an optimization problem for findingthe current distribution in inner and outer field-generating surfacessatisfying all constraints and having the minimum magnetic stored energyand/or dissipation. These inner and outer field-generating surfaces willthen be generated into current loop designs. Next in step 306 theoptimization problem is solved using magnetic design software. Andfinally in step 308 the continuous stream functions obtained whichdefine the current distributions in the inner and outer field-generatingsurfaces which were determined as a result of the optimization patternare converted into a pattern of discreet current loops, and connectingthese discreet current loops in series to define a gradient coil design.The current distributions in the inner and outer field-generatingsurfaces are defined by the continuous stream function.

Actively shielding the gradient coil only at the locations ofsuperconducting magnet sections, letting the external fieldunconstrained in the regions between these sections, may improve theefficiency of the gradient system, compared to a fully actively shieldedconfiguration, by a factor 2. This enables building a wide bore systemusing a magnet similar to the cryo-free magnet now being developed for anarrow-bore low-cost scanner. Actively shielded gradient coils areinherently inefficient because most of the field energy is located inthe region between the primary and shield windings.

Allowing the field to expand into the external space makes the coil moreefficient but would result in unwanted magnetic coupling with the mainmagnet. Such unwanted coupling may be avoided by using the coil designdisclosed here.

The external field of the gradient coil is shielded only there wheresuperconducting coil sections of the main magnet are located and is freeto expand into the magnet environment in the spaces between magnetwindings.

The radiation screen can be made transparent to AC magnetic fields, byproviding a large number of axial interruptions in the conductingcylinder. The magnet bore can be made of non-conducting material;optionally the gradient coil can serve as the cryostat bore tube. Theimprovement in efficiency can be used to increase the patient bore. Agradient coil compatible with a patient space of 700 mm diameter andwith a stored energy of less than 5 joule at 10 mT/m can be fittedinside a magnet with a coil ID of 886 mm. A magnet of this size would bejust large enough for a narrow bore scanner when a conventional gradientcoil is used. The external field shaping concept can but need not becombined with non-cylindrical shapes of the inner gradient windings. Theconcept can also be used to reduce the diameter and cost of the magnetfor a narrow bore scanner. When applied to z-gradients, the externalfield must be shaped such that no net flux is coupled into the mainmagnet sections.

If the superconducting magnet is not enclosed in a conventional tankfilled with liquid helium and if there is no other electricallyconducting cylinder between the magnet sections, then the external fieldof the gradient coil need only be constrained at the location of themagnet windings. With suitable gradient design tools, it is possible toexploit this and generate gradient designs where the field of thegradient “bubbles out” between these magnet sections. The result of thesimulations described here is that the efficiency of the gradient coilcan be improved by a factor of 1.5-2. This gain can be used to decreasethe distance between primary and shield coils of the gradient coil,which either reduces size and cost of the magnet on the outside of thegradient coil or increase the patient bore at constant magnet diameter.

The sections of the superconducting magnet may be re-optimized tomaximize the space for the external gradient field to pass through, butthis is a matter of straightforward magnet design.

The gradient coil to be used in this concept has a different shape butcan be built using the same technology as conventional actively shieldedgradient coils now used in the product.

1. Introduction

The gradient system in a whole-body MR system is one of the main costdrivers in the system. It is squeezed between the patient space with theRF system and the main field magnet and optimizing any of these (largerbore, smaller, cheaper magnet) drives the cost and complexity of thegradient system up. One of the reasons why conventional activelyshielded gradient coils takes up so much space is the fact that currentdesigns seek to reduce the external magnetic field of the gradient coilfrom interacting with the superconducting magnet and its cryostat. Anyfield escaping from the outer surface of the gradient tube will induceeddy-currents in the electrically conducting cylinders of the cryostatand field modulation in the superconducting magnet windings will causeincreased dissipation in the cold mass. If it would be possible to leaveaway the active shielding layer; the primary gradient windings could belocated near the inner bore of the cryostat. Such an unshielded gradientcoil would, however, create far too much field in the cryogenicenvironment of a conventional helium bath-cooled MRI magnet.1.1. Objective

Simulations have been performed to find out whether it makes sense torelax the external field requirements of a gradient coil, still keepingit decoupled from the windings of the superconducting main field magnet,but allowing field in the spaces between these winding sections. For thesolutions found, the key properties of the gradient coil (stored energy,dissipation) were calculated and compared to the same parameters of anequivalent gradient coil designed using conventional methods andrequirements.

2. Simulation Model and Method

The gradient coil was modeled in the usual way as two concentriccylindrical meshes. In all simulations, the field quality inside thecoil was defined by a set of constraint points on an ellipsoidal volumeextending 450 mm in radial direction and 360 mm in z-direction, with amaximum deviation from a linear gradient field of 0.3 mT at a gradientstrength of 10 mT/m. The reference case of complete stray-fieldcompensation was modeled in the conventional way, by limiting theinduced current in a secondary surface at a radius of 10 mm outside theouter gradient layer. For the relaxed stray field model, the sections ofthe superconducting magnet were modeled as toroidal secondary surfacescovering the outer boundary of the superconductor bundles. (Wallthickness 1 mm, resistivity 10⁻⁹ Ohm·m). The magnet was a modifiedversion of a small-bore 1.5T magnet design (warm bore ˜820 mm), withsmaller length/thickness ratio of the coil sections. This changeincreases the peak field on conductor but does not have a large impacton the amount of conductor. For each value of the outer radius of thegradient coil a new magnet was generated. An example of the modelgeometry is shown in FIG. 4.

FIG. 4 shows a mesh 400 which defines the magnetic gradient coil in twolayers and the smaller cylinders labeled 402 represent meshes definedfor the coils of the superconducting magnet.

The objective to keep the external field of the gradient coil out of thevolumes occupied by the superconducting wire could have been modeled asconstraints limiting the normal component of the field on all boundarysurfaces of the toroidal wire volumes. The simpler way to achieve theobjective is to limit the dissipation in the surfaces bounding thecoils. This was done by putting a heavy weight factor on thecontribution of this dissipation to the cost function and by biasing theoverall optimization towards minimizing the dissipation.

FIG. 5 shows an example of a typical example solution of the relaxedexternal fields. 500 shows the design of an inner conductive layer. Thecircular lines indicate current loops and 502 represents the outerconductive layer of the magnetic gradient coil. The cylinders labeled108 are again portions of the superconducting magnet and the solution ascalculated by the magnetic design software is shown above the innerconductive layer 500, the outer conductive layer 502 and thesuperconducting coils 108. It can be seen that there is an expansion ofthe magnetic field 504 between adjacent conductors of thesuperconducting magnet.

A typical example of a solution with the relaxed external fieldconstraints in shown in FIG. 5. The current pattern in the shield coilstill resembles that of a classical active shield coil, but the shape ofthe windings has become more irregular. The field plot shown in FIG. 5clearly demonstrates that the field is bulges out into the spacesbetween the coil sections. As will be shown in more detail in the nextsection, relaxing the external field strongly reduces both the storedenergy and the dissipation in the gradient coil. Typically, both numbersare at least a factor 1.5 lower for the relaxed shielding case.

With the constraints used in the simulation, the typical dissipation inthe magnet sections is approximately 0.5 W or less (for 10 mT/m RMS,100% duty cycle). The boundaries of the coil were modeled as 1 mm thickcopper sheet at helium temperature, resistivity 10⁻⁹ Ohm-m. If thesections of the real magnet are enclosed in such copper liners (in sucha way that no closed conducting loops are formed) the dissipation can bekept within the cooling range of the refrigerator.

3. Results of Parametric Studies

In principle, the gain in efficiency resulting from relaxing theexternal field constraints can be used in two ways. Either the externaldimensions of the system can be reduced at constant size of the patientbore or the patient bore can be increased at constant size of themagnet. Both pathways were explored in this study.

3.1. Compressing the Magnet at 600 mm Patient Bore

For a gradient coil with an inner radius of 325 mm the outer radius ofthe coil was varied between 360 and 425 mm. The results are shown inFIG. 6 and FIG. 7. Act shield designates a conventional fully shieldedgradient coil, act shaped is short for a coil where the field is shapedsuch as to stay clear of the magnet windings.

FIG. 6 shows a graph of the stored energy 600 versus the coil outerradius 602. The coil outer radius is given in terms of meters and thestored energy is in arbitrary units. This is shown for fully shieldedgradient coil 604 that is conventionally shielded versus an example of afield shaped gradient coil. The field shaped gradient coil shapes isoperable to generate a magnetic field that avoids the superconductivecoils. It can be seen that the stored energy is significantly lower forthe field shaping.

FIG. 7 shows the same plot for the conditions of FIG. 6 except thedissipation of energy 702 is shown instead of the stored energy. Againit can be seen that the field-shaped gradient coil 606 has significantlyless energy dissipation than the fully shielded coil 604. FIGS. 6 and 7illustrate that examples of the magnetic gradient coil as describedherein can greatly reduce the amount of stored energy and dissipationwhen compared to conventional magnetic gradient coils.

FIGS. 6 and 7 clearly show that active field shaping instead of fullyshielding results in a significantly more efficient coil than a fullyactively shielded coil. The difference in stored energy is approximatelya factor 1.5 over the entire range of coil dimensions; the advantage ofimperfect shielding increases somewhat on reducing the outer radius ofthe coil. FIG. 3 shows that it is possible to reduce the diameter of themagnet by about 100 mm until the stored energy of the relaxed gradientdesign reaches the level of the fully shielded coil.

The dissipation is also reduced by a factor 1.5 by relaxing the externalshielding requirements on the coil design; the ratio increases as theouter radius of the coil is reduced. The dissipation values are thosecalculated by the modeler for current patterns for a copper thickness of2 mm.

3.2. Magnet for Partly Shielded Gradients

In all subsequent simulations, the coil inner radius was set to 443 mmand the coil configuration was kept unchanged. The magnet used in thesesimulations is shown in FIG. 8. It is a classical 6+2 configuration witha good homogeneity within a deformed ellipsoid extending 430 mm intransverse direction and 360 mm in axial direction.

FIG. 8 shows the magnetic field contours 800 in a model of a magnetdefined by the superconducting coils 108. FIG. 8 shows the main fieldmagnet with 443 mm inner coil radius, contours are 3, 10, 30, and 100uT. The gradient simulations showed, that it the exact location of thesuperconducting sections does not have a significant effect on thefeasibility of gradient designs making use of the empty space betweenthese coils. More important for the success of the approach is the shapeof the magnet windings: the best gradient designs are obtained when themagnet windings have nearly square cross section (except for the endcoils, these can be of high aspect ratio).

3.3. Wider Bore Gradient Coils Inside a Small-Bore Magnet

A series of gradient coils was simulated with increasing inner radiusand a constant outer radius of 415 mm. The smallest of these coilscorresponds to a 600 mm patient bore, the largest would nearly be largeenough for a wide bore system. For comparison, the equivalent fullyactively shielded gradients were also simulated (with an outer radius of415 mm). Similar to what was seen before, the potential advantage ofactive field shaping increases when the gradient coil is compressed.

The graph shows, that even for the largest coil in the series theefficiency of the coil would be similar to conventional designs. Thepredicted dissipation is higher, but would be still at an acceptablelevel. FIG. 9 and FIG. 10 also contain a few points for coils with anouter radius of 425 mm. Bringing the outer coil close to the magnet onlyslightly improves the stored energy but has a significant effect on thedissipation.

FIGS. 9 and 10 show the results of some simulations for a series ofgradient coils which were simulated with increasing inner radius and aconstant outer radius of 415 mm. FIG. 9 shows the stored energy 602versus the inner radius 900. The inner radius is given in meters and thestored energy 602 is in arbitrary units. FIG. 10 shows the dissipationversus the inner radius. The inner radius is labeled 900 again and isgiven in meters and the dissipation 702 is given in arbitrary units.These plots are shown for constant outer radius of 415 mm 902 and isalso shown in several cases for an outer radius of 425 mm also. Thecurves labeled 906 shows the result for the complete active shielding.The dot labeled 908 shows the inner coil with a recess. FIG. 9 shows thestored energy of partly and fully shielded gradient coils vs. innerradius, for outer radius 415 (425) mm and magnet radius 443 mm. Dot:inner coil with recess, 360/380 mm, OR 425 mm, recess length 600 mm.

FIG. 10 shows the Dissipation in partly and fully shielded gradientscoil vs. inner radius, same geometries as FIG. 9.

In all simulations the predicted dissipation in the main magnet sectionsremained below 0.5 W for 1 mm cold copper. The predicted dissipation inthe magnet did not change much when the outer radius of the gradientsystem was increased.

3.4. Field and Conductor Patterns

A typical example of the external field of the partially shieldedgradient coil is shown in FIG. 2. The field is nicely constrained in theregions of the main magnet coils and expands outward in the spacebetween the coils. FIG. 11, below, is another visualization of the samecoil geometry and field data, after discretization into 250 ampereconductors. In oblique planes through the z-axis of the system the fieldis also small at the magnet coils and much larger in between. Here thefield also has an azimuthal component; the flux is partly carried aroundthrough the z=0 plane, partly through the x=0 plane. An importantconsequence of this azimuthally oriented return field is that theradiation screen shall not form closed loops when going once around thesystem via the loop bore—end_flange1—outer_rollup—end_flange2. Thisrequirement also holds for internal wiring inside the cold mass.

FIG. 11 shows the visualization of a coil geometry as was shown in FIG.5. In FIG. 11, the Gradient inner radius is 370 mm, the outer radius is425 mm, magnet inner radius 443 mm.

The flat conductor patterns corresponding to this coil are shown inFIGS. 12 and 13. Even with a very small number of turns the externalfield accurately follows the contour found in the continuous currentdistribution. With this selection of operating current, the inductanceof this coil would be 155 microHenry.

FIG. 12 shows the conductor pattern 1200 of the conductive layer 500.The conductive layer 500 is cylindrical so FIG. 12 shows the conductorpattern 1200 laid out flat. Similarly FIG. 13 shows the conductorpattern 1300 of the outer conductive layer 502.

3.5. Recessed Primary Coil Option

Further increase of the inner diameter of the coil would drive thedissipation probably higher than desired. One way to further increasethe patient bore would be to provide a recess in the primary coil and tolocate the RF coil in this recess. FIG. 9 and FIG. 10 also contain onedata point for a gradient coil with an inner recess, which would allow afurther increase in patient bore space. The recess had a depth of 20 mmand a total length of 600 mm (allowing a long and good quality RF coil).In general, the active field shaping concept can be combined with allother methods to improve the gradient efficiency (includingnon-cylindrical and/or a-symmetric cross-section of the inner bore).

3.6. Z-Gradient Coil

The z-gradient coil is generally much more efficient than the transversecoils, so it is less urgent to attempt to optimize its efficiency. A fewsample simulations were done to see whether incomplete shielding is alsoeffective for this channel. Relaxing the external field for these coilshas to be done with care, because the z-gradient can become inductivelycoupled to sections of the main magnet. If the gradient coil causes anet flux in any of the sections of the main coil, this can lead to largeinduced voltages if the gradient coil is switched. Therefore, inaddition to minimizing the field at the main magnet windings, az-gradient system with relaxed external field shielding also has tosatisfy the requirements that the mutual inductance with each magnetsection be close to zero. The partly shielded z-gradient coil wasgenerated on the same mesh as that of the transverse coil with innerradius 370 mm, outer radius 425 mm and magnet radius 435 mm. Theresulting coil is shown in FIG. 10. The contour lines show the modulusof the field, with contour step size 0.1 mT (@10 mT/m). This coil wouldhave a stored energy of 2.8 J (90 microHenry at 10 mT/m@250 A). Thepredicted dissipation in the magnet is less than 100 mW. This is goodenough to have confidence that a decent z-gradient can be incorporatedinto the concept, without introducing any additional technologicaldifficulties.

4. Considerations for Practical Implementation

The main differences will occur in the superconducting magnet and in theinterface between the magnet and the gradient coil. To facilitate makingthe concept work, the magnet sections are preferably short inz-direction, providing maximum free space in between. The supportingstructure of the magnet preferably does not support eddy currents.Because there will always be some residual external gradient field, themagnet windings need a copper or aluminum liner (which will probably bethere anyway in order to keep the magnet sections at operatingtemperature. The conducting liners should preferably fully enclose thecoils.

The radiation screen and the magnet bore tube have to be made such thatthey do not support eddy currents. For the room temperature bore, thiscan be done using available technology: all magnets we had prior to 1989had glass-fiber reinforced plastic bore tubes. One option for this wallwould be to make use of the outer cylinder of the gradient coil as theinner wall of the outer vacuum container. The parametric analysis of theprevious chapter shows, that the penalty for a separate gradient coilinside a thicker walled bore tube is not dramatic. A beneficial sideeffect of a non-conducting magnet bore is, that no eddy currents can beinduced in it. This is likely to reduce the acoustic noise generated bythe system.

A slit radiation screen with good thermal conductivity in axialdirection but high thermal and electrical resistance in azimuthaldirection could be made from two slit patterns of 2-3 mm thick plate,glued together, in such a way that the conducting elements are orientedpredominantly in axial direction. The conducting strips are then eitherconnected alternatingly to one end flange or connected to both when theyhave an interruption somewhere along the length.

While the invention has been illustrated and described in detail in thedrawings and foregoing description, such illustration and descriptionare to be considered illustrative or exemplary and not restrictive; theinvention is not limited to the disclosed embodiments.

Other variations to the disclosed embodiments can be understood andeffected by those skilled in the art in practicing the claimedinvention, from a study of the drawings, the disclosure, and theappended claims. In the claims, the word “comprising” does not excludeother elements or steps, and the indefinite article “a” or “an” does notexclude a plurality. A single processor or other unit may fulfill thefunctions of several items recited in the claims. The mere fact thatcertain measures are recited in mutually different dependent claims doesnot indicate that a combination of these measured cannot be used toadvantage. A computer program may be stored/distributed on a suitablemedium, such as an optical storage medium or a solid-state mediumsupplied together with or as part of other hardware, but may also bedistributed in other forms, such as via the Internet or other wired orwireless telecommunication systems. Any reference signs in the claimsshould not be construed as limiting the scope.

LIST OF REFERENCE NUMERALS

-   100 magnetic resonance imaging system-   102 magnet-   104 bore of magnet-   106 cryostat-   108 superconducting coil-   110 magnetic gradient coil-   112 magnetic gradient coil power supply-   114 imaging zone-   116 radio frequency coil-   118 radio frequency transceiver-   120 subject-   122 subject support-   130 cylindrical axis of symmetry-   132 length parallel to axis 130-   134 outer surface of magnetic gradient coil-   150 computer system-   152 hardware interface-   154 processor-   156 user interface-   158 computer storage-   160 computer memory-   162 pulse sequence-   164 magnetic resonance data-   166 magnetic resonance image-   168 control module-   170 image reconstruction module-   200 magnetic resonance imaging system-   202 vacuum-   204 radiation shield-   206 warm bore-   208 inner cylinder of radiation shield-   400 mesh defining gradient coil-   402 mesh defining super-conducting coil-   500 inner conductive layer-   502 outer conductive layer-   504 expansion of magnetic field-   600 outer radius of coils [m]-   602 stored energy [A.U.]-   604 fully shielded gradient coil-   606 shield shaped to avoid superconducting coils-   702 dissipation of energy [A.U.]-   800 magnetic field contours-   900 inner radius [m]-   1200 conductor pattern-   1300 conductor pattern

The invention claimed is:
 1. A magnet assembly for a magnetic resonanceimaging system comprising: a superconducting magnet with multiplesuperconducting coils, and wherein there is a superconducting coilselected from the multiple superconducting coils centered coaxiallyabout each reduced field region; and an actively shielded gradient coiloperable for generating a magnetic field, wherein the magnetic field hasa cylindrical axis of symmetry, wherein the gradient coil has a lengthparallel with the cylindrical axis of symmetry, wherein the magneticgradient coil has an outer surface, wherein the magnetic field comprisesan external magnetic field outside of the outer surface, and wherein theexternal magnetic field has at least four reduced field regions alongthe length where the modulus of the magnetic field is less than theaverage of the modulus of the magnetic field along the length.
 2. Themagnet assembly of claim 1, wherein the linear extent of each of thereduced field regions as measured along the length is at least 10% ofthe distance between two adjacent reduced field regions.
 3. The magnetassembly of claim 1, further including at least one selected from thegroup consisting of: the modulus of the external magnetic field withinany one of the reduced field regions is at least a factor of 2.5 timessmaller than the average of the modulus of the magnetic field along thelength; the modulus of the external magnetic field within any one of thereduced field regions is at least a factor of 5 times smaller than theaverage of the modulus of the magnetic field along the length; themodulus of the external magnetic field within any one of the reducedfield regions is at least a factor of 10 times smaller than the averageof the modulus of the magnetic field along the length; and the modulusof the external magnetic field within any one of the reduced fieldregions is at least a factor of 20 times smaller than the average of themodulus of the magnetic field along the length.
 4. The magnet assemblyof claim 1, wherein the gradient coil has an inner conductive layer andan outer conductive layer, wherein the inner conductive layer comprisesa first set of discrete current loops connected in series, and whereinthe outer conductive layer comprises a second set of discrete currentloops connected in series, and wherein the first set is connected inseries to the second set.
 5. The magnet assembly of claim 1, wherein thegradient coil comprises three orthogonal gradient coils, wherein thereduced field region of the orthogonal gradient coils coincide topositions of at least some of the coils of the superconducting magnet.6. The magnet assembly of claim 1, wherein the magnet comprises a warmbore tube and a radiation screen, and wherein the radiation screencomprises an inner cylinder of the radiation screen between the warmbore tube and the multiple superconducting coils.
 7. The magneticassembly of claim 6, wherein the warm bore tube and the inner cylinderof the radiation screen are electrically non-conducting or have a higherelectrical impedance in the azimuthal direction than in the axialdirection of the cylinder and/or wherein the inner cylinder is formedfrom a dielectric.
 8. The magnetic assembly of claim 7, wherein theinner cylinder of the radiation screen is formed from a conductivematerial with slots operable to block eddy currents generated by theexternal magnetic field of the magnetic gradient coil or the innercylinder of the radiation screen is formed from a dielectric.
 9. Themagnet assembly of claim 6, wherein the warm bore tube comprises themagnetic gradient coil.
 10. The magnet assembly of claim 1, wherein themagnet is a cryogen free magnet.
 11. The magnet assembly of claim 1,wherein the magnetic gradient coil is operable for producing theexternal magnetic field such that it expands between eachsuperconducting coil.
 12. A magnetic resonance imaging system comprisinga magnet assembly according to claim
 1. 13. A magnetic gradient coil fora magnetic resonance imaging system, wherein the magnetic gradient coilis actively shielded, wherein the magnetic gradient coil is operable forgenerating a magnetic field, wherein the magnetic field has acylindrical axis of symmetry, wherein the gradient coil has a lengthparallel with the cylindrical axis of symmetry, wherein the magneticgradient coil has an outer surface, wherein the magnetic field comprisesan external magnetic field outside of the outer surface, and wherein theexternal magnetic field has at least four reduced field regions alongthe length where the modulus of the magnetic field is less than theaverage of the modulus of the magnetic field along the length, whereinthe linear extent of each of the reduced field regions as measured alongthe length is at least 10% of the distance between two adjacent reducedfield regions.
 14. A method of designing a magnetic gradient coil for amagnetic resonance imaging system using magnetic design software, themethod comprising: A. defining cylindrical surfaces corresponding to aninner conductive layer and an outer conductor layer of the gradient coiland defining conducting flange regions connecting the inner conductivelayer and the outer conductive layer; B. defining constraints enforcinga magnetic gradient field with a predetermined linearity with an imagingvolume inside the gradient coil, wherein the magnetic gradient coil hasan outer surface; C. defining constraints limiting an outer magneticfield surrounding the outer surface such that that there are at leastfour reduced field regions corresponding to the location of asuperconducting coil of a superconducting magnet and/or definingsurfaces acting as passive conducting rings at the locations of thereduce field regions and defining constraints for the current induced inthese rings or the dissipation caused by these induced currents, whereinsteps A, B, and C define an optimization problem for finding currentdistributions of the inner conductive layer and the outer conductivelayer; D. solving the optimization problem to calculate a continuousstream function, wherein the continuous stream function is descriptiveof the solution to the optimization problem; and E. converting thecontinuous stream function obtained as the result of the optimizationinto a pattern of discrete current loops, wherein the stream andconnecting these discrete current loops in series to define a gradientcoil design.
 15. The method of claim 14, further comprisingmanufacturing the gradient coil according to the gradient coil design.16. A method of magnetic resonance imaging comprising: generating amagnetic field longitudinally through an imaging zone with asuperconducting magnet which includes at least four superconducting coilrings separated by annular regions; with an actively shielded gradientcoil disposed between the superconducting magnet and the imaging zone,generating gradient magnetic fields across the magnetic field in theimaging zone and a leakage magnetic field outside the actively shieldedgradient coil, wherein the leakage magnetic field has at least fourreduced field regions, each reduced field region being longitudinallyadjacent one of the superconducting coil rings, wherein a modulus of theleakage magnetic field within any one of the reduced field regions is atleast a factor of 2.5 smaller than an average of a modulus of theleakage magnetic field along a length of the superconducting magnet;receive magnetic resonance signals from a subject in the imaging zonewith at least one radiofrequency coil; with one or more processors,reconstruct the magnetic resonance signals into an image; display theimage on a display device.